Scintillation crystals are commonly used in non-invasive medical diagnostic techniques that utilize radiation emitting materials. These crystals are noted for their ability to emit pulses of visible light when ionizing radiation, such as gamma radiation, passes therethrough and interacts with atomic nuclei in the crystal. The pulses of emitted light (photons) are then detected by a photodetector device such as a photomultiplier tube (PMT) or a semiconductor photodiode (SPD). The effectiveness of the detector in diagnostic procedures depends on the ability to see and quantify the crystal light flashes with high spatial, spectral, and temporal precision. This, in turn, is dependent on brightness and rapidity of the generated and recorded flash, which are functions of the type and geometry of the scintillation crystal.
Position emission tomography (PET) is an example radiation imaging technology that provides in-vivo, functional information about the molecular biochemistry of a given radio-labeled compound (a.k.a. tracer) introduced into a live subject. The radio-label is a positron emitter. The tomographic imaging is possible through detection and localization of the many associated highly energetic photons emitted. The sensor is typically an array of scintillation crystals. There are essentially three stages of the photon sensing. The entering photon is first absorbed in the scintillation crystal. The crystal gives off a flash of light. This light is collected by a photodetector, which subsequently detects and converts the light into electric charge that is amplified. The result is a robust electric signal with an amplitude that represents the energy of the incoming photon, a location that indicates where the energetic photon came from within the imaging subject, and time stamp that signifies when the event occurred. For high spatial resolution imaging, which will allow one to see very minute structures, PET relies on very accurate localization of the energetic photon emissions. This means the scintillation detector must have very fine position resolution of the entering photons. However, to efficiently absorb the incoming photons, the crystal must also be relatively thick. Efficient absorption of incoming photons is important to allow for high count sensitivity, which translates into good image quality.
In typical commercial devices, there has been a compromise between detector spatial resolution, detection efficiency, and light collection into the photodetector. High light collection is important for creation of robust electronic detector signals for high sensitivity. High light collection is also important for good energy resolution for good photon Compton scatter and random coincidence background rejection capabilities. Scatter and random photon coincidences produces loss of image contrast and should be rejected as much as possible. Efficient light collection is also important for good event timing resolution that further helps to reject random coincidence background.
In typical commercial PET devices spatial resolution is determined by the dimensions of the individual scintillation crystals. Typical commercial nuclear medicine cameras utilize a two-dimensional, discrete or pseudo-discrete array of long, narrow scintillation crystals, which are coupled at a small end to photomultipliers (PMTs) with the opposite small end directed toward the high energy photon source. Positional information is encoded in the crystals, as each crystal indicates a unique X-Y position. The crystals are made long for high gamma-ray stopping power and narrow for high spatial resolution. An intermediate optical coupling medium is necessary in these designs at the scintillation crystal/PMT interface. Crystal surfaces in these designs are treated and coated with reflectors to preferentially direct light through internal reflections into the PMT located at one small end of the scintillation crystal array. Detection sensitivity in the conventional discrete-crystal designs is hampered by coupling of the scintillation crystals to the PMTs. Another type of commercial device used in Nuclear Medicine utilizes a scintillation crystal sheet, with the radiation incident on the large surface of the crystal sheet. Detection by a PMT is on an opposite large surface of the crystal sheet. This configuration is commonly known as a scintillation camera.
The typical commercial PET devices collect emitted light photons at the opposite end from the end at which the radiation of interest is received in the long, narrow crystals. The front as well as the side crystal surfaces include reflective coatings to help alleviate loss of intensity in the scintillation light resulting from the received radiation. However, for very fine crystals in high resolution PET systems only a small fraction of the scintillation light produced in the crystal reaches the photodetector. This light loss problem associated with standard photodetector readout at the end of the scintillating crystal worsens as the crystal is made narrower and longer or has untreated surfaces. This light loss problem together with the low quantum efficiency of the PMT photocathode for detecting the scintillation light produced limits the count efficiency and signal-to-noise ratio of both the crystal decoding scheme used to position and time an annihilation photon event, and the energy (spectral) resolution required to reduce gamma-ray scatter and random coincidence background rates. Good scatter and random photon reduction is an important factor for improving image contrast between true structures of interest and the background present in the resulting PET images.
An additional related problem associated with the conventional end readout is that the light collection efficiency depends on the location within the crystal where light was created and thus, where the radiation interacted. This factor further degrades the energy resolution. Also, there is roughly a 10–15% light loss at the interface between the crystal and PMT due to index of refraction mismatches, further degrading the signal to noise ratio. In a particular commercial high resolution design, additional fiber optics are included between the crystals and the PMT, which further degrades the available light signal.
Extracting a high fraction of the available scintillation light from the ends of long and narrow crystals proves to be very difficult due to a poor aspect ratio for light collection. The result is lower signal to noise ratios (S/N), relatively small pulse heights (reduced sensitivity), and inadequate energy resolution (reduced Compton scatter and random photon rejection capabilities). This low light extraction also contributes to non-optimized coincidence time and spatial resolution. To facilitate light collection, the crystal sides must be highly polished, which significantly increases complexity and costs.
The state of the art was advanced by the invention described in U.S. Pat. No. 6,114,703 (incorporated by reference herein) to Levin et al. The '703 patent provided an efficient method and devices for collection, and made the large surfaces of long and narrow scintillation crystals available for detection. The '703 patent disclosed methods and devices that replaced the bulky and expensive PMTs by utilizing semiconductor photodetectors, applying such semiconductor photodiodes directly to surfaces of the scintillation crystals, including at least one large surface of the scintillation crystal. The device of the '703 patent improved the amount of light measured from a scintillation event, while maintaining high spatial resolution offered by long and narrow scintillation crystals. The '703 patent also improved upon the single sheet style conventional devices that receive radiation perpendicular to the large face of the crystal sheet by eliminating the coupling losses associated with the optical interfaces between the crystal and PMT and replacing the PMT of the conventional devices with directly deposited semiconductor photodiodes.
An overriding goal in radiation imaging is to obtain reconstructed images of very high spatial resolution. Spatial resolution improvements in reconstructed images have come most often from reductions in the size and increases in the number of scintillation crystals. Detection sensitivity, though, is another limiting factor. The '703 patent was directed to improvements in the detection sensitivity. To maintain high detection sensitivity and good image quality, the challenges are to develop a finely pixellated scintillation crystal array with both high detection efficiency and high light collection. High detection efficiency means the crystals must be relatively long, tightly packed, and cover a relatively large axial field-of-view (FOV).
A difficulty with designs having small scintillation crystals for high resolution is that manufacturing is a significant challenge. It is costly and complex to handle many minute crystal elements and align them with corresponding photodetector elements. Slight misalignments might reduce light collection efficiency. A shortcoming with conventional crystal sheet devices for PET is that the sheet must be thin so that it produces a relatively narrow beam of light onto the photodetector plane. Thus, crystal sheet detectors that have been used in PET suffer from low efficiency for stopping the high energy photons.